The human respiratory tract has an extensive blood supply and its total surface area is about 75-140 square meters in adults, features which make it attractive as a route for administering medicaments (drugs). Local administration of drugs intended to target pulmonary sites of action, as in asthma or chronic obstructive pulmonary disease (COPD), is the traditional application for drug delivery by the respiratory route. However, systemic administration of drugs through the lungs is also attractive for substances that undergo metabolic breakdown in the gastrointestinal tract and are unsuitable for oral administration. Large molecules, such as peptides and proteins, that cannot pass the absorbing membrane after oral administration (insulin as one example) are potential substances for which pulmonary route of administration may be preferable to other routes. Additionally, drugs that undergo extensive ‘first-pass’ metabolism in the liver may be better to administer via the pulmonary route. The lungs have inherent anatomical and physiological advantages for administration of such drugs. The lungs are robust organs that provide a large (75-140 m2) and well-perfused alveolar surface for absorption with a thin alveolar-capillary barrier and only modest regional variation in the amount of mucus production. Most of the lungs' surface area resides in the alveolated regions of the deep lung, which also contains a rich capillary network to assure efficient gas exchange during the breathing process. There is a lack of mucociliary clearance at the alveolar region so that drug that is deposited in the deep lung is not likely to be expelled and is likely to be absorbed into the bloodstream. The lungs also lack certain peptidases that are present in the gastrointestinal tract, so that bioavailability of peptide or protein drugs may approach that for parenteral routes of administration.
As with other drug delivery routes that are alternatives to oral administration, there are potential challenges to be overcome with drugs being delivered through the lungs. This is because of the number of factors that can strongly influence the efficacy of drug delivery via the pulmonary system, including: particle size, particle density, particle surface morphology, particle charge, hygroscopicity of the particles, rate of dissolution of particles at the site of deposition, the patient's breathing pattern (especially the inspiratory flow rate and the extent of inhalation; slow and deep inhalation tend to increase alveolar deposition), patient comorbidities (such as interstitial lung disease or airflow obstruction), losses within the device or the environment, as well as other factors. Among these factors, particle size distribution is arguably the most critical. The optimal particle size (aerodynamic diameter) for deep (alveolar) pulmonary deposition of particles is around 1-3 μm in diameter. Larger particles become lodged in the upper respiratory tract. Particles smaller than about 0.5 μm tend to be exhaled and not deposited in the respiratory tract at all. In this connection, it is important to note that the size distribution of inhaled aerosolized drug particles is not a constant, as-manufactured size distribution. The sizes of particles in the aerosol plume change continuously during the course of the patient's inhaled ‘puff’ (inspiratory maneuver), due to a variety of physical processes such as agglomeration (coagulation) through collisions with other inhaled particles, condensation of water vapor on the particles, evaporation in the case of liquid particles, and so forth.
The main challenge in the design of formulations to be delivered to the lungs is the incorporation of the drug in particles with an appropriate aerodynamic size distribution. Past research efforts investigated the deposition profiles for an inhaled nebulizer cloud as predicted from model which assume oral breathing and healthy subjects. For example, some studies determined that oropharyngeal deposition decreases with decreasing median particle diameter, falling from 60% of the inhaled dose at 10 μm to almost 0% below 1 μm. Central airway deposition peaks at 5-7 μm and peripheral airway deposition at 2-3 μm. However, it must be considered that usually the patient's respiratory system status in disease is an important factor that affects directly the behavior of the particles once inhaled. As previously mentioned, particles smaller than 0.5 μm are frequently not deposited in the respiratory tract but are exhaled.
The aerodynamic diameter refers not only to the geometric particle diameter, but also to the particle mass density. Larger but porous particles have been also proposed as an option for reducing inertial impaction of particles onto surfaces in the throat, thereby improving the chance that the particles will reach the alveolar region. However, large, porous particles have large surface area and, in general, are more susceptible to condensation of water vapor from the humid airstream during their progress through the respiratory flow path. In the case of highly hygroscopic particles, this can not only lead to increases in particle size but also increases in particle density.
As the aerosol plume passes from the inhaler device into the patient's airways, the particle size may change by evaporation of the volatile components (solvents) on the surface of inhaled liquid droplets, by condensation of water vapor from the humid airways upon the inhaled droplets or powder particles, and by agglomeration of particles that collide with one another during transit. The transit time during the inspiratory maneuver by which the aerosol plume is inhaled into the patient's respiratory tract has a duration of hundreds of milliseconds or longer. The particle size distribution changes in the particle plume during this time may be substantial. In many cases, the mean particle size may increase by a factor of two or more, by processes of particle-particle collisions and agglomeration (coagulation). The larger particles undergo greater inertial impaction on mouth and oropharynx and other structures of the upper respiratory tract or are precipitated and settle within the inhaler device itself. Therefore, the development of effective pulmonary delivery systems that can reach the alveoli is still a challenge. Furthermore, the interaction of the aerosol plume may be significantly influenced by respiratory physiology, inspiratory airflow parameters, and attributes of a patient's upper airway anatomy. In particular, aerosolization mechanisms that involve vibratory elements that emit acoustic frequencies at audible and ultrasonic frequencies may either accelerate or inhibit particle agglomeration, depending on these attributes.
The generation of clouds of droplets by means of acoustic waves was first reported in 1927. Two different mechanisms explain atomization, capillary waves and, in the case of ultrasonic frequencies, cavitation. The Kelvin equation for capillary waves is described as:
  λ  =            (                        2          ⁢          πσ                          ρ          ⁢                                          ⁢                      f            2                              )              1      /      3      
where:
λ=Wavelength of surface waves at the air/liquid interface
σ=Surface tension coefficient
ρ=Liquid density
f=Frequency of the surface waves
After further investigations, the Kelvin equation was later modified and the following expression derived:
  λ  =            (                        8          ⁢          πσ                          ρ          ⁢                                          ⁢                      F            2                              )              1      /      3      
where F is the forcing sound frequency or frequency of the acoustic signal. These theoretical bases were already established when in 1927 the possibility of atomizing liquid by exciting them with ultrasonic waves was described. Decades later, it was determined that the expression relating the wavelength to droplet size D could be through an empirical constant, which was experimentally reported to be 0.34.
  D  =      0.34    ⁢                  (                              8            ⁢            πσ                                ρ            ⁢                                                  ⁢                          F              2                                      )                    1        /        3            
Based on this principle, nebulizer devices currently in the market operate at relatively high frequencies (usually 20 to 200 KHz) aiming to impart energy to the liquid efficiently to generate small droplet size distributions, while formulations to be nebulized under this principle should have higher densities and lower surface tension values in order to achieve lower droplet diameters. Likewise, vibratory dry power inhaler (DPI) devices tend also to operate at relatively high frequencies (usually greater than 2 KHz) aiming to transfer energy efficiently to the powder to cause deagglomeration and nearly complete dispersal of the solid powder into a plume having a small particle size distribution capable of achieving extensive alveolar deposition.
Some dry powder inhalers and nebulizer devices employ vibratory means to disperse an amount of a drug powder into an aerosol that is carried on the stream of inspired air as an aerosol ‘plume’ (referred to herein as “vibratory inhalers”). With vibratory inhalers, acoustical vibrations created by the dispersal elements have a frequency spectrum that generally interacts with the acoustic properties of the patient's mouth, pharynx, and upper airway. The transit time for air between the inner aspect of the teeth to the upper aspect of the vocal folds is sufficiently long that significant coagulation of powder particles can occur en route. If the frequency spectrum of the vibratory inhaler device is matched to the acoustic properties of the patient's upper respiratory flow-path, the acoustic vibrations emitted by the vibratory dispersal element may substantially inhibit the undesired coagulation and precipitation and/or impaction upon the surfaces of these upper structures, permitting more of the powdered pharmaceutical formulation to enter the deep lung and proceed to the preferred sites of deposition. If on the other hand the frequency spectrum of the vibratory inhaler device is ill-matched to the acoustic properties of the patient's flow-path, then coagulation and precipitation or impaction may be significantly worse, resulting in non-efficacy of the inhaled drug due to failure of the device to deliver the drug to the proper site of action.
The fluid mechanical ‘resistance’ of a vibratory inhaler device has been evaluated as it relates to the ability of a patient to mount an adequate inspiratory effort to achieve adequate air velocities through the device. However, little, if any, attention has been paid to the acoustic properties of vibratory inhaler dispersal mechanisms and their relationship to vocal tract acoustic spectral properties and to changes in the ‘fine particle fraction’ (FPF) and decreases in particle count concentration as the inhaled aerosol plume progresses into the respiratory tract.
Conventional single and multiple dose dry powder inhaler devices use: (a) individual pre-measured doses in blisters containing the drug, which can be inserted into the device prior to dispensing; or (b) bulk powder reservoirs which are configured to administer successive quantities of the drug to the patient via a dispensing chamber which dispenses the proper dose. In operation, vibratory inhaler devices strive to administer a uniform aerosol dispersion amount in a desired physical form of the dry powder (such as a particulate size) into a patient's airway and direct it to a desired deposit site(s).
A number of obstacles can undesirably impact the performance of dry powder inhalers. For example, the small size of the inhalable particles in the dry powder drug mixture can subject them to forces of agglomeration and/or cohesion (certain types of dry powders are susceptible to agglomeration, which is typically caused by particles of the drug adhering together), which can result in poor flow and non-uniform dispersion. In addition, as noted above, many dry powder formulations employ larger excipient particles to promote flow properties of the drug. However, separation of the drug from the excipient, as well as the presence of agglomeration, can require additional inspiratory effort, which, again, can impact the stable dispersion of the powder within the air stream of the patient. Unstable dispersions may inhibit the drug from reaching its preferred deposit/destination site and can prematurely deposit undue amounts of the drug elsewhere. Similar obstacles also affect active vibratory nebulizer devices. Further, some vibratory inhalers can retain a significant amount of the drug within the device, which can be especially problematic over time.
Some inhalation devices have attempted to resolve problems attendant with conventional ‘passive’ inhalers. For example, U.S. Pat. No. 5,655,523 to Hodson et al. proposes a dry powder inhalation device which has a deagglomeration/aerosolization plunger rod or biased hammer and solenoid, and U.S. Pat. No. 3,948,264 to Wilke et al. proposes the use of a battery-powered electromechanical element to vibrate the capsule to effectuate the release of the powder contained therein. These devices propose to facilitate the release of the dry powder by the use of energy input independent of patient respiratory effort. U.S. Pat. No. 6,029,663 to Eisele et al. proposes a dry powder inhaler delivery system with a rotatable carrier disk having a blister shell sealed by a shear layer that uses an actuator that tears away the shear layer to release the powder drug contents. The device also proposes a mouthpiece cover that is attached to a bottom portion of the inhaler. U.S. Pat. No. 5,533,502 to Piper proposes a powder inhaler using the patient's inspiratory effort to generating a respirable aerosol and also includes a rotatable cartridge holding the depressed wells or blisters defining the drug reservoirs. A spring-loaded carriage compresses the blister against conduits with sharp edges that puncture the blister to release the drug that is then entrained in air drawn in from the air inlet conduit so that aerosolized drug is emitted from the aerosol outlet conduit and into the patient. The published U.S. Pat. Appl No. 20070209661 of Smyth et al. also proposes a passive aeroelastic mechanism to vibrate and disperse powder using patient inspiratory efforts in a manner that is tolerant of a wide range of inspired air velocities and flow-rates. The contents of all of these patents and applications are hereby incorporated by reference as if stated in full herein. Hickey et al., in U.S. patent application Ser. No. 10/434,009 and PCT Patent Publication No. WO 01/68169A1 and related U.S. National Stage patent application Ser. No. 10/204,609, have proposed a DPI system to ‘actively’ facilitate the dispersion and release of dry powder drug formulations during inhalation using vibrating piezoelectric elements which may promote or increase the quantity of fine particle fraction particles dispersed or emitted from the device over conventional DPI systems. The contents of these documents are hereby incorporated by reference as if recited in full herein. Each of these active and passive inhaler devices imparts energy to the drug in order to deagglomerate it and aerosolize it, and does so with acoustic vibrations of the airstream, either directly (as the primary means of aerosolization) or indirectly (secondary to the motions of the elements involved in generation of the aerosol).